1. Field of the Invention
The present invention relates to a method for the determination and use of a standard operational value for the delay time of a radiographic system used in medical radiographic applications. More in particular the method relates to a technique to determine the delay time or start-time of a generator, used in combination with a digital radiographic detector, a so-called FPD, Flat Panel Detector, and to use same as standard operational value for radiographic exposures.
2. Description of the Related Art
It is known that radiographic illumination has important applications in medical imaging, whereby the medical advantages for the patient largely exceed the small risk of damage resulting from such radiographic illumination. The formation of images is a result of the fact that radiographic illumination, depending on the energy, passes through most soft tissues, but does not pass the harder calcium-containing tissue. So, as an example, bones will bar most radiographic illumination whereas cartilage will bar such illumination to a much lesser extent. As a result, since more than a century, the human skeleton can easily be visualised by radiographic illumination. Radiographic illumination is also applied in medicine in radio therapeutic applications. For the application of the present invention however radiographic illumination is used for medical imaging applications, more in particular for medical diagnosis.
Radiographic illumination generally is generated in an X-ray tube as ‘Bremsstrahlung’. An example of such an x-ray tube is shown in FIG. 2.
This radiation is generated when accelerated electrons impinge on a target, mostly made of tungsten of another hard material such as molybdenum with a melting point above 2 000 degrees Celsius. The electrons are accelerated under vacuum (10-5 Pascal) by use of an electric current. A tension difference (anode tension) generates such field between the cathode K and the anode A. Electrons are liberated from the Cathode K by heating same, for example by a resistance wire of filament, whereupon an incandescent tension is applied, causing an incandescent current that generates a local heating.
At around 2 700 K electrons are generated around the tungsten filament that can escape from such material at these elevated temperatures. As such electrons are negatively charged, they are accelerated by a strong electric field (30-150 kV depending on the application) over a distance of a few millimetre from the negative cathode to the positive anode. The maximal energy of the radiographic illumination generated in this way is proportional to the electrical voltage applied; therefore is usually is expressed as kilo-electron volt. The intensity is depending on the electrical current that is generated. This is expressed in mA (mill amperes).
When generating such intense radiographic illumination, the target, the anode, is heated intensively. As a result, radiographic tubes are provided with cooling arrangements (for example water cooling and/or are characterised by a high rotational speed of the target (rotating-anode).
The anode mostly is made of metal. In the anode material the electrons are intensively slowed down, producing a radiographic illumination having an energy comprised between O and the total voltage of the electric field. This illumination is called Bremsstrahlung. On top hereof a high number of such electrons will be slowed down by collision with electrons in the anode material and will ionise atoms by liberating electrons from the inner shells. When such electrons fall back to such inner shells, so-called characteristic radiation is generated, depending on the kind of metal the anode is made of. In case of a copper-anode this radiation is around 8 keV, whereas in case of molybdenum this radiation amounts to approx. 18 keV. The total charge of a radiation tube is a few kilowatts, the surface whereupon the electrons impinge is between 0.5 and 10 square millimetre.
For use in medical imaging applications, two kinds of radiation tubes are used, differing as regards the construction of the anode. All tubes consist of a glass tube wherein the all components are present under high vacuum. One type of radiation tube comprises a fixed anode, whereas the other comprises a rotating anode. Because all tubes and casings are fully closed, it is not possible to use a cooling medium from outside, as was the case in the past. The only possibility is the discharging of the energy generated by radiation. To this end, the tube is surrounded by an amount of air-free high-quality oil. A switch is foreseen that will automatically interrupt the current in case of expansion of the oil by heating; in this way the tube is protected against overheating.
This is the original model, whereby the anode usually is made of copper, characterised by an excellent heat conductivity.
This model is mostly used in apparatuses with a limited power as used for example in dentist applications, in portable and mobile units.
To achieve an improved discharge of the heat, the rotating anode concept has been developed. In this concept, use is made of a massive disc of tungsten or an alloy of tungsten and rhenium. The place where the electrons impinge is not limited to approximately 1 square centimetre, but consists of a circle over the disc surface, the so-called line focus. Also a second incandescent filament can be mounted in the cathode, that focuses on a smaller of larger surface on the anode. A small target surface (focus) is characterised by less scattering and thus less geometrical unsharpness. For small objects (hands, feet, small joints) one usually chooses the smallest possible focus for a maximal rendering of details.
The anode is formed by the disc, the support and the anode body that functions as the rotor of an electromotor. At the outside magnets are mounted (stator) that enable the anode to quickly turn around. Depending on the tube-type the rotating velocity is situated between 4000 and 9000 revolutions per minute. The angle of the anode is usually situated between 10° and 20°, which is much smaller than in case a fixed anode is used. Tubes having a rotating anode usually have much more components that fixed-anode tubes and the steering thereof requires much more electronic circuitry. Consequently these models are much more expensive compared to the more simple models. The efficiency of tubes with a rotating anode is however much higher and the application of this type of tubes practically has no limits.
Preferred embodiments of the present invention as described hereinafter relate to radiation tubes with a rotating anode; these types of tubes are most common nowadays, in particular for general purpose radiography (genrad), as well as for mammographic applications (mammo).
As is known to those skilled in the art, in case of a radiographic illumination with a DR Panel, the generator should first receive a signal that the anode should be brought to speed, and the filament or incandescent wire of the cathode should be heated to a red/white state.
After a certain amount of time which is required for the above—this is the so-called generator delay time or start-up time—the expose or illumination button can be activated, whereby the generator is brought under high tension.
In case the operator activates both buttons simultaneously, or in case of a combined prep-expose button, activates the button at full, a pre-determined fixed delay time will cause the high tension to occur only after the anode is brought to speed and the filament or incandescent wire is heated to the red/white state.
This start-up or delay time of the radiographic generator should be known.
These times differ for the various types of radiographic generators that are on the market.
This problem in particular arises when an existing radiographic exposure unit in a hospital was used in combination with radiographic detectors such as film or stimulable phosphors, and now should be used in combination with fully digital radiographic detectors, such as flat panel detectors.
This is the so-called retrofit situation, known to the person skilled in the art of medical radiography.
One of the crucial differences between the use of radiographic films or stimulable phosphors, as contrary to fully digitized panels, is that films and stimulable phosphors are always ‘ready’ to be exposed.
The only limitation for a film is that it should not be exposed and/or developed in an earlier stage, and in case of a stimulable phosphor screen, that it has been erased after a prior exposure.
Provided these conditions are met, both radiographic media are always apt to be used in a radiographic exposure.
The radiographic workflow in case a fully digitized radiographic panel is used, on the contrary, is rather different. The reason for this is that in most cases a radiographic digital panel should first be reset. This resetting is known to the person skilled in the art, and is described amongst others in the Handbook of Medical Imaging, Vol. 1, Chapter 4: Flat Panel Imagers for digital radiography, (ed. R. V. Matter et al., SPIE Press, Bellingham, 2000.)
The resetting of a radiographic digital panel, and the problems caused hereby in case of a change-over of an existing radiographic exposure unit, previously used in combination with film or stimulable phosphor plates, to a unit based on the use of digital radiographic detectors, is published in a great number of earlier patent specifications, amongst others in EP 2 209 422.
The global aim in case of a radiographic exposure is that the integration time of the panel or the digital radiographic detector overlaps with the exposure-time of the generator. More in particular, the integration time should somewhat exceed the exposure time to be sure no radiographic diagnostic information is lost.
To this end it is essential to know the applicable generator delay time.
Indeed, suppose the start-up of the generator would coincide with the integration time of the direct radiographic panel, then it is not excluded that the integration time has expired at the moment the effective exposure starts. Such a procedure evidently would not lead to an image useful for medical diagnosis.
In practice, a radiographic exposure can take place according to two different ways. First the radiographic operator activates the prep button of the apparatus, e.g. on the retrofit box.
The patient to be radiographed is then requested to keep still and not to breathe temporarily (in case of a chest exposure). As set forth supra, the generator then receives a signal to bring the anode to speed and to heat the filament.
In case the operator waits sufficiently long for the activation of the expose button, until the generator's rotating anode is effectively up to speed (the so-called generator delay time), no problem arises.
When the exposure button is activated by the operator, a signal is sent to the DR panel to reset same, and once this is done, over the retrofit box a signal is sent to the generator to start the exposure.
The generator is ready to perform same because its generator delay time has expired, in other words, because the generator has had enough time to prepare the exposure.
The problem arises when—as is often the case in practical circumstances—the operator activates, e.g. by pressing, the prep and exposure buttons simultaneously, or in case the operator pushes through (in case the prep and exposure buttons are integrated in one and the same button or device, that can be activated as well in part as in full). In such a case, the problem can arise that the generator receives an expose signal at a time the generator is not ready for this, in other words, because its anode is not yet brought to speed, and/or the filament wire is not sufficiently heated.
In such a preferred embodiment, the generator will necessarily wait to commence the exposure until the time the rotating anode is up to speed, but in the meantime the panel is ready to receive the radiographic illumination, differently phrased, the panel is integrating same, but there does not exist a meaningful radiographic signal to integrate. In the worst case scenario, the generator starts to perform exposure at a time the panel again is closed, because for example its integration time has expired.
In such scenario the above situation often leads to a so-called retake, which means that the patient will be exposed again. This evidently should be avoided, given the inherent harmful effect of any radiographic exposure on the health condition of patients.